1. Field of the Invention
The present invention generally relates to a nuclear medical diagnostic device (an emission computed tomography (ECT) device), in which an radioactive agent is applied to a test subject, and a pair of γ-rays discharged by positron radioactive isotopes (RIs) accumulated in a target portion of the test subject is measured concurrently, so as to obtain a tomogram of the target portion. In particular, the present invention relates to a technique of counting γ-rays at the same time.
2. Description of Related Art
A positron emission tomography (PET) device is taken as an example to illustrate a nuclear medical diagnostic device, that is, an ECT device. The PET device is formed in the following manner. Opposite γ-ray detectors are used to detect two γ-rays at an angle of approximately 180° discharged from a target portion of a test subject. When the γ-rays are detected (counted) at the same time, a tomogram of the detected body is formed again. Further, some of the γ-ray detectors used to count the γ-rays at the same time in the PET device are formed by scintillators and photomultipliers. The scintillators then emit light after the γ-rays discharged from the test subject are incident thereon, and the photomultipliers convert the light emitted by the scintillators to an electric signal.
In principle and in most cases, the γ-rays discharged from a visual field centre are obliquely incident on the scintillators of the γ-ray detector D as shown in FIG. 16. When the divided scintillator is not present in the γ-ray incident direction, the γ-ray is detected in both the correct positions and incorrect positions. That is, the visual difference error becomes larger from the visual field centre to the peripheral parts, such that the tomogram obtained by the PET device becomes inaccurate.
Therefore, as shown in FIG. 17, a γ-ray detector is provided. In the γ-ray detector, the scintillators are divided (optically combined) into scintillators having the luminescence pulses in the γ-ray incident direction with different attenuation times. For example, when a γ-ray detector MD is used, in which the scintillators are divided into a scintillator array with a shorter γ-ray attenuation time on a γ-ray incident side and a scintillators array with a longer γ-ray attenuation time on a photomultiplier side, the positions of the discharged γ-rays can still be detected accurately even if the γ-rays are obliquely incident on the scintillators of the γ-ray detector MD. A more accurate tomogram is resulted, and improvement is achieved (for example, please refer to Patent Reference 1 and 2).
Further, the specific mechanism for detecting the γ-ray position with respect to the scintillator array with the shorter attenuation time and the scintillator array with the longer attenuation time, stacked in the γ-ray incident direction, includes the following mechanisms: an adding mechanism for converting an electric signal output from a light receiving element, that is, converting an analog signal SF (the signal of the scintillator array with the shorter attenuation time) or SR (the signal of the scintillator array with the longer attenuation time) to a digital signal by using an A/D converter as shown in FIG. 18, and adding sequentially the digital signals converted by the A/D converter as shown in FIG. 19; an identification value calculating mechanism for calculating an identification value representing a value AT1/AT2 or BT1/BT2 obtained by dividing an intermediate added value AT1 or BT1 by a total added value AT2 or BT2, in which the intermediate added value AT1 or BT1 is a value obtained by adding the digital signals during a period starting from the point at which the luminescence of the luminescence pulse is emitted from the scintillator block to an intermediate point of the period at which the luminescence ends, that is, to an intermediate moment in the adding mechanism, and the total added value AT2 or BT2 is a value obtained by adding the digital signals during a period starting from the point at which luminescence of the luminescence pulse is emitted from the scintillator block to point at which the luminescence ends, in the adding mechanism; a mechanism for deciding a medium value K according to a maximum value and a minimum value in the identifying values calculated by the identification value calculating mechanism; and a determination mechanism for determining whether an identification value calculated by the identification value calculating mechanism is greater or smaller than the medium value K.
Moreover, the existing nuclear medical diagnostic device determines the parameters used for judging in the following manner. That is, in the case of a two-stage scintillator detector 112, for example, having scintillator arrays of a two-stage structure as shown in FIG. 20, parameters T1, T2, and K required by the scintillator array identification mechanism are decided in the following manner. In an inspection stage of the γ-ray detector single unit, the two-stage scintillator detector 112 disposed in a dark box 115 inputs the parameter of an initial value to a processing circuit for inspection, starts to irradiate γ-ray on a front surface 110 of the scintillator array, and then calculates a signal count N1 and a signal count N2 through a determination calculation, in which the signal count N1 is determined as the count of signals from the front surface 110 of the scintillator array, and the signal count N2 is determined as the count of signals from a rear surface 111 of the scintillator array. Thereafter, as shown in FIG. 21, the γ-ray is only irradiated on the rear surface 111 of the scintillator array, and a signal count N2′ and a signal count N1′ are calculated through a determination calculation, in which the signal count N2′ is determined as the count of signals from the rear surface 111 of the scintillator array, and the signal count N1′ is determined as the count of signals from the front surface 110 of the scintillator array. Further, as shown in FIG. 22, when a ray source is not used, and on a background of a natural radioactive ray 116, a signal count N1b and a signal count N2b are calculated through a determination calculation, in which the signal count N1b is determined as the count of signals from the front surface 110 of the scintillator array, and the signal count N2b is determined as the count of signals from the rear surface 111 of the scintillator array. Then, (N1−N1b)/(N2−N2b) and (N2′−N2b)/(N1′−N1b) are defined. When (N1−N1b)/(N2−N2b) and (N2′−N2b)/(N1′−N1b) are equal and are the maximum, the parameters are defined as the optimal parameters. Further, a lead collimator 13 and a Ri ray source 114 are required to ensure that the γ-ray is irradiated on only any one of the scintillator arrays. The parameters decided in the above manner are pre-input to a processing circuit for the device during the stage in which the γ-ray detector single unit is installed in the actual PET device.
Patent Reference 1: Japanese Patent Publication No. H06-337289 (Page 2 to 3 and FIG. 1)
Patent Reference 2: Japanese Patent Publication No. 2000-56023 (Page 2 to 3 and FIG. 1)
However, the existing nuclear medical diagnostic device has the following problem. In the case of the two-stage scintillator detector 112, for example, having the scintillator arrays of the two-stage structure as shown in FIG. 20, the parameters T1, T2, and K required by the scintillator array identification mechanism are decided by the processing circuit for inspection during the inspection stage of the γ-ray detector 12 single unit, and these parameters are applied to the processing circuit for the device during the stage in which the γ-ray detector single unit is installed in the actual PET device. No matter the processing circuit for inspection and the processing circuit for the device are manufactured based on the same specification, temperature characteristics of a gain amplifier or other elements may be slightly different, so as to generate individual differences. Therefore, the optimal values of the parameters are not necessarily consistent, such that it is impossible to separate upper and lower parts, resulting in adverse impact on the image quality.
In another aspect, if it is intended to decide the parameters during the stage in which the γ-ray detector single unit is installed in the actual PET device, a large lead calibration jig and Ri ray source matching with the PET device are required, so the operation is quite complicated.